1. Field of the Invention
The present invention relates to an improved porous polymer useful for various applications in industry, including the medical industry, for example, as a biological prosthesis and particularly useful in vascular surgery. The porous polymer can be made by use of a new gel enhanced phase separation technique, which, among other advantages, permits enhanced shape-making capability.
2. Discussion of Related Art
The present invention encompassing polymer engineering and processing came about from efforts to improve existing properties of porous polymers, including medical devices and prostheses and, in particular, medical devices (e.g., vascular grafts). Accordingly, a review of the vascular graft art is appropriate.
The search for the ideal blood vessel substitute has to date focused on biological tissues and synthetics. Despite intensive efforts to improve the nature of blood vessel substitutes many problems remain, such as increasing failure rate with decreasing caliber of the blood vessel substitute, a high failure rate when infection occurs, and aneurysm formation. The major need for vascular grafts is for adequate supply of blood to organs and tissues whose blood vessels are inadequate either through defects, trauma or diseases. Vascular grafts are also needed to provide access to the bloodstream for individuals undergoing hemodialysis. The three major types of vascular grafts are peripheral, arterial-to-venous access, and endovascular.
Peripheral grafts are those used in the neck and extremities, with the most common being used in the leg. This results in supply problems being some intermediate and most small diameter arteries are replaced or bypassed using an autologous saphenous vein, the long vein extending down the inside of the leg, with a secondary source being the radial veins of the arms. In a given patient, suitable veins may be absent, diseased or too small to be used, and removal of the vein is an additional surgical procedure that carries attendant risk.
Additionally, arterial-to-venous access grafts are used to access the circulatory system during hemodialysis. Vascular grafts used in connection with hemodialysis are attached to an artery at one end and sewn to a vein at the other. Two large needles are inserted into the graft. One needle removes the blood where it flows through an artificial kidney machine and is then returned to the body via the second needle. Normal kidney function is destroyed by several acute and chronic diseases, including diabetes and hypertension. Patients suffering from kidney failure are maintained by dialysis three times a week for approximately four hours per session. Due to the constant punishment these grafts undergo, there is a high occurrence of thrombosis, bleeding, infections, and pseudoaneurysm.
Endovascular grafts are used to reline diseased or damaged arteries, particularly those in which aneurysms have formed, in a less invasive manner than standard vascular surgical procedures. Various surgical techniques and materials have been developed to replace and repair blood vessels. Ideally, the thickness of the prosthesis is minimized, so that it can be delivered to the implantation site using a percutaneous procedure, typically catheterization and kept in place utilizing stents. Problems associated with this type of implantation include thrombosis, infection and new aneurysm formation at the location of the stent.
Initially, autografts were used to restore continuity; however, limited supply and inadequate sizes forced the use of allografts from both donor and umbilical cord harvest such as that described in U.S. Pat. No. 3,974,526. Development of aneurysms and arteriosclerosis as well as the fear of disease transmission necessitated the search for a better substitute.
Artificial vascular grafts are well known in the art. See for example U.S. Pat. No. 5,747,128; U.S. Pat. No. 5,716,395; U.S. Pat. No. 5,700,287; U.S. Pat. No. 5,609,624; U.S. Pat. No. 5,246,452 and U.S. Pat. No. 4,955,899. Development of two different fibrous and pliable synthetic plastic cloths revolutionized vascular reconstructive surgeries. Whenever suitable autograft was not available woven grafts of polyethylene terephthalate (Dacron®) and drawn out polytetrafluoroethylene (Teflon®) fibrils as defined in U.S. Pat. Nos. 3,953,566; 4,187,390 and 4,482,516 were used. Even though these products were widely used they did have many drawbacks including infection, clot formation, occlusions and the inability to be used in grafts smaller than 6 mm inside diameter due to clotting. Additionally, the graft had to be porous enough so that tissue ingrowth could occur, yet have a tight enough weave to the fibers so that hemorrhage would not occur. This made it necessary to pre-clot these grafts prior to use. Recently, vascular prostheses have been coated with bioabsorbable substances such as collagen, albumin, or gelatin during manufacture instead of preclotting at surgery. For purposes of this patent disclosure, the term “bioabsorbable” will be considered to be substantially equivalent to “bioresorbable”, “bioerodable”, “absorbable” and “resorbable”.
Compliance problems with woven polyethylene terephthalate and drawn out polytetrafluoroethylene prompted interest in thermoplastic elastomers for use as blood conduits. Medical grade polyurethane (PU) copolymers are an important member of the thermoplastic elastomer family. PU's are generally composed of short, alternating polydisperse blocks of soft and hard segment units. The soft segment is typically a polyester, polyether or a polyalkyldoil (e.g., polytetramethylene oxide). The hard segment is formed by polymerization of either an aliphatic or aromatic diisocyanate with chain extender (diamine or glycol). The resulting product containing the urethane or urea linkage is copolymerized with the soft segment to produce a variety of polyurethane formulations. PU's have been tested as blood conduits for over 30 years. Medical grade PU's, in general, have material properties that make it an excellent biomaterial for the manufacture of vascular grafts as compared to other commercial plastics. These properties include excellent tensile strength, flexibility, toughness, resistance to degradation and fatigue, as well as biocompatiblity. Unfortunately, despite these positive qualities, it became clear in the early 1980s that conventional ether-based polyurethane elastomers presented long-term biostabilty issues as well as some concern over potential carcinogenic degradation products. Further, in contrast to excellent performance in animal trials, clinically disappointing results with PU-based grafts diminished the attractiveness of the material for this application.
Recent developments in new generation polyurethanes, however, have made this biomaterial, once again, a promising choice for a successful long-term vascular prosthesis. Specifically, the new generation of polyurethanes solved the biostability problems but still provide clinically disappointing results. Poor performance is largely due to limitations of current manufacturing techniques that create a random or non-optimal fibrous structure for cell attachment using crude precipitation and/or filament manufacturing techniques. (See, for example, U.S. Pat. Nos. 4,173,689; 4,474,630; 5,132,066; 5,163,951; 5,756,035; 5,549,860; 5,863,627 & International Patent Publication WO 00/30564)
Nonwoven or non-fibrous polyurethane vascular grafts have also been produced, and various techniques have been disclosed for swelling and/or gelling polyurethane polymers.
U.S. Pat. No. 4,171,390 to Hilterhaus et al. discloses a process for preparing a filter material that can be used, for example, for filtering air or other gases, for filtering gases from high viscosity solutions, or for preparing partially permeable packaging materials. A first solution containing an isocyanate adduct dissolved in a highly polar organic solvent is admixed into a second solution containing a highly polar organic solvent and a hydrazine hydrate or the like. The first solution is admixed into the second solution over an extended period of time, during which time the viscosity of the admixture increases as the hydrazine (or the like) component reacts with the isocyanate to produce a polyurethane. The first solution is added up to the point of instantaneous gelling. The final admixture is coated onto a textile reinforcing material, and the coated material is placed in a water bath to coagulate the polyurethane. The resulting structure features a thin, poreless skin that must be removed, for example, by abrasion, if the structure is to be useful as a filter.
U.S. Pat. No. 4,731,073 to Robinson discloses an arterial graft prosthesis comprises a first interior zone of a solid, segmented polyether-polyurethane material surrounded by a second zone of a porous, segmented polyether-polyurethane, and usually also a third zone surrounding the second zone and having a composition similar to the first zone. The zones are produced from the interior to the exterior zone by sequentially dipping a mandrel into the appropriate polymeric solution. The porous zone is prepared by adding particulates such as sodium chloride and/or sodium bicarbonate to the polymer resin to form a slurry. Once all of the zones have been formed on the mandrel, the coatings are dried, and then contacted to a water bath to remove the salt or bicarbonate particles.
U.S. Pat. No. 5,462,704 to Chen et al. discloses a method for making a porous polyurethane vascular graft prosthesis that comprises coating a solvent type polyurethane resin over the outer surface of a cylindrical mandrel, then within 30 seconds of coating, placing the coated mandrel in a static coagulant for 2-12 hours to form a porous polyurethane tubing, and then placing the mandrel and surrounding tubing in a swelling agent for 5-60 minutes. After removing the tubing from the mandrel, the tubing is rinsed in a solution containing at least 80 weight percent ethanol for 5-120 minutes, followed by drying. The coagulant consists of water, ethanol and optionally, an aprotic solvent. The swelling agent consists of at least 90 percent ethanol. The resulting vascular graft prosthesis features an area porosity of 15-50 percent and a pore size of 1-30 micrometers.
U.S. Pat. No. 5,977,223 to Ryan et al. discloses a technique for producing thin-walled elastomeric articles such as gloves and condoms. The method entails dipping a mandrel modeling the shape to be formed into a coagulant solution, then dipping the coagulant coated mandrel into an aqueous phase polyurethane dispersion, removing the mandrel from the dispersion, leaching out any residual coagulant or uncoagulated polymer, and finally curing the formed elastomeric article. When the polyurethane dispersion comprises by weight about 1 to 30 parts per hundred of a plasticizer based on the dry polyurethane weight, the dispersed polyurethane particles swell. Thus, if the dispersion featured polyurethane particles having a mean size between 0.5 and 1.0 micrometer in the unplasticized condition, they might be between 1.5 and 3.0 micrometers in the plasticized condition. The inventors discovered that such swollen polyurethane particles produce a superior product, whereas in an unplasticized condition, particles of such a size (1.5-3.0 micrometers) impede uniform drying because of the large interstitial space between particles. Preferred coagulants are ionic coagulants such as quaternary ammonium salts; preferred plasticizers are the phthalate plasticizers.
U.S. Pat. No. 3,492,154 to Einstman et al. discloses a technique for making a microporous sheet, useful as an artificial leather material. Einstman dissolves a polymer such as polyurethane in a solvent to make a solution, coats the solution onto a porous substrate, and then plunges the coated substrate into a liquid that coagulates the polymer. The solvent preferably is miscible in the coagulant. The coagulant is a non-solvent such as water, methanol, benzene or chloroform. Important to a high quality product is rapid coagulation of solid polymer from solution, and Einstman accomplishes this by dropping the temperature of the coated substrate by at least 10 C during coagulation, and also by adding a quantity of the non-solvent to the polymer solution prior to quenching. The quantity is not enough to cause coagulation at the higher temperature, but rather is about 70 to 98% of this quantity.
U.S. Pat. No. 3,553,008 to Reischl et al. is also concerned with producing leather-like sheets. To carry out his process, Reischl dissolves polyurethane in a solvent. As with Einstman, Reischl adds a quantity of non-solvent to his solution, preferably not enough to cause gel formation, but preferably at least 60% of this quantity. The non-solvent should be miscible with the solvent. The solution is then applied to a substrate, which can be either porous or non-porous. The solvent and non-solvent are then evaporated, leaving behind solid, microporous polyurethane.
U.S. Pat. No. 5,077,049 to Dunn discloses injecting a polymeric solution into the human body. The solvent for the polymer is water-miscible, so upon contact with aqueous saline solution (e.g., body intercellular fluid), the latter extracts the solvent. As is disclosed in other prior art documents, this contact with aqueous solution coagulates the polymer, thereby forming a solid or gelatinous implant. Dunn discloses a long list of candidate solvents, as well as a long list of polymers. Dunn goes on to explain that not all of his listed solvents will dissolve every polymer listed (col. 5, lines 52-58). There is some overlap among the liquids disclosed by Dunn, and those of the present invention. However, there is no disclosure in Dunn of mixing or combining solvents, let alone a disclosure or suggestion of applying liquids in sequential fashion to gel a polymer solution. Dunn uses water for his gelation step, water being a well-known non-solvent (his claim reads “water-coaguable).
In each instance, there are severe shape-making limitations, e.g., the known non-fibrous methods appear to be limited to working with a relatively low viscosity liquid that can be coated onto a surface, or into which a shape-forming mandrel can be dipped. It would be desirable if the polymer could be rendered in the form of a gel because a gel, inter alia, can be molded, such as being extruded. In other words, the gel can be plastically shaped and can retain its molded shape without reverting to its original shape. Usually the molded shape is preserved so that the shaped polymer retains the new shape and will return to the new shape if deformed, provided that the elastic limit is not exceeded. Further, most of the above-discussed non-fibrous art results in a product that features a non-porous layer at least at some location in the product. Thus, the prior art does not seem to appreciate the desirability of a prosthesis such as a vascular graft containing channels or porosity extending continuously from the exterior surface to the luminal surface of the graft.
One of the reasons for failure of vascular grafts is due to the formation of acute, spontaneous thrombosis, and chronic intimal hyperplasia. Thrombosis is initiated by platelets reacting with any non-endothelialized foreign substance, initiating a platelet agglomeration or plug. This plug continues to grow, resulting in occlusion of the graft. If the graft is not immediately occluded the plug functions as a cell matrix increasing the potential for rapid smooth muscle cell hyperplasia. Under normal circumstances, platelets circulate through the vascular system in a non-adherent state. The endothelial cells lining the vascular system accomplish this. These cells have several factors that contribute to their non-thrombogenic properties. These factors include, but are not limited to, negative surface charge, the heparin sulfate in their glycocalyx, the production and release of prostacylin, adenosine diphosphate, endothelium derived relaxing factor, and thrombomodulin. Thus, adherence of a thin layer of endothelial cells to the vascular prosthetic results in enhanced healing times and reduced failure rates of the graft.
Other reasons for artificial graft failure are neointima sloughing due to poor attachment and aneurysm formation resulting from compliance mismatch of the new graft material to the existing vascular system. It is important to know that materials with different mechanical properties, when joined together and placed in cyclic stress systems, exhibit different extensibilities. This mismatch may increase stress at the anastomotic site, as well as create flow disturbances and turbulence. Additionally, poor attachment geometry can lead to the problematic results above, due to flow disturbances and turbulence. For example, the harvesting of autograft veins typically causes a surgeon to use a graft of non-optimal diameter or length. A graft diameter mismatch, of perhaps 60% or more, causes a drastic reduction in flow diameter. Such flow disturbances may lead to para-anastomotic intimal hyperplasia, anastomotic aneurysms, and the acceleration of downstream atherosclerotic change.
Finally, artificial graft failures have been linked to leaking of blood through the device. Pre-clotting and the addition of short-lived bioabsorbable substances such as collagen, gelatin and albumin can prevent this as well as provide a matrix for host cell migration into the prosthesis. One problem with this approach is that the same open fibrous weave that permits blood leaking also allows the viscous bioabsorbable substances and clotted blood to accumulate on the luminal surface and easily detach resulting in complications (e.g., emboli) downstream from the device.